X-ray detection systems for such applications as computerized tomography commonly employ combinations of scintillating crystals and photodiodes. For example, a CAT scanner system operates by taking multiple, cross-sectional, X-ray slices from different angles within a single plane passing through a body. An X-ray source and an array of detectors are placed on opposite sides of an annular gantry, which rotates in the selected plane around the body. Signals generated by the detector array are digitized and mathematically processed to create a cross-sectional image of the body.
In a scintillating-photodiode X-ray detection system, the incident X-rays are absorbed by a scintillating crystal and converted into visible light. That visible light is then absorbed into a silicon photodiode, which converts the light into electron-hole pairs that diffuse from the P-N junction and thereby could generate a current flow. Because the current flow is typically of very small magnitude, it is common to use an amplification means to amplify the photodiode signal and convert it into a voltage. The output of such a scintillating-photodiode-preamplifier system is a voltage that is proportional in magnitude to the incident X-ray flux on the scintillating crystal. Systems of this type are described in a chapter by Promod Hague entitled "Scintillator crystal-photodiode array detectors" appearing in Thomas H. Newton and D. Gordon Potts (eds.), "Technical Aspects of Computed Tomography," vol. 5 at 4127-4132 (1981), which chapter is incorporated herein by reference.
In a typical photodiode construction, the silicon wafer is appropriately doped so as to create a narrow P-type zone or region adjoining a first face of the wafer and a narrow N-type zone adjoining a second, opposite face, the P and N zones being separated by an almost intrinsic region in the interior portion of the wafer. For example, it is conventional to create P-type zones using a boron dopant and N-type zones using a phosphorus dopant. The photodiode is mounted on a substrate, for example, along the N-type face, and a scintillating crystal is mounted along the P-type face using silicon grease or other optically-transparent epoxy as a coupling medium between the adjoining scintillating crystal and photodiode surfaces. Electrical terminals are connected respectively to the P and N zones to collect the current flow generated by the photodiode.
The conventional photodiode construction described above is generally well suited for detecting light in the infrared range because these light photons generate electron-hole pairs at internal locations relatively distant from the surface. Thus, accurate readings of infrared radiation require that electrical charge be collected over much of the internal volume of the silicon. By contrast, it is well known that X-ray scintillating crystals typically produce blue light photons having a wavelength of about one-half that of infrared radiation. Unlike infrared light, the blue light photons penetrate on the order of only several microns into the silicon. Accordingly, for such X-ray detection applications, applicants have found that it is only necessary to collect the light-generated electrical charge from the surface of the photodiode and from those internal regions immediately adjacent to the photodiode surface.
At the same time, however, because of the very low level of the electrical signals associated with X-ray detection applications, the familiar problems of "noise," "electrical cross-talk," "optical cross-talk" and "electrical response" become highly significant. "Electrical cross-talk" in this usage refers to a phenomenon that can occur in an array of multiple, adjacent scintillating crystal-photodiode pairs and their respective associated electrodes. If an electrical signal generated by an X-ray flux incoming to a first scintillating crystal-photodiode pair is accidentally "collected" at the electrodes associated with an adjacent scintillating crystal-photodiode pair, the result is an erroneous detection reading.
"Electrical response" here refers to the problem that a thicker electrical medium slows, reduces in magnitude, and may distort electrical signals passing through that medium. If a conventional X-ray photodiode construction is used for blue light photons, for example, an electrical signal generated in a 1-3 micron photodiode "depletion zone" adjacent to the scintillating crystal-photodiode interface would have to travel without recombination through as much as 300 microns of the silicon in order to reach the associated cathode terminal, thereby resulting in relatively slow and possibly inaccurate detection readings. It is, therefore, desirable to reduce the length of the electrical pathway that a generated electrical charge must traverse in order to reach the associated electrical terminal.
"Optical cross-talk" as used herein refers to the problem that a portion of the light photons generated in the scintillating crystal of a first scintillating crystal-photodiode pair may pass into a second, adjacent photodiode. Similarly, a light photon from a first scintillating crystal may be reflected off a metallized electrical contact, deflected into an adjacent crystal, and then directed into the photodiode associated with that adjacent crystal. The result in either case is the generation of an electrical charge in the second, adjacent photodiode instead of in the photodiode of the first scintillating crystal-photodiode pair, again resulting in a detection error. A related "noise" problem occurs when a portion of an X-ray flux to a scintillating crystal passes into an electrically-active zone of the silicon and generates charges by direct ionization of the silicon. Again, the result of the "noise" phenomenon could be an erroneous detection reading.
In addition to overcoming or minimizing the foregoing problems with conventional X-ray photodiodes and photodiode arrays, it is commercially desirable to produce X-ray photodiodes that are easier, faster and less expensive to make, and that have a higher degree of uniformity. These and other problems with and limitations of the prior art photodiode designs are largely overcome with the coplanar X-ray photodiodes, coplanar photodiode arrays, and coplanar photodiode X-ray detection systems of this invention.